In at least some computed tomograph (CT) imaging system configurations, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the “imaging plane”. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal spot. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator adjacent the collimator, and photodiodes adjacent the scintillator.
The scintillator material converts X-ray radiation into nearly monochromatic (in the case of HiLite scintillators, commercially made by General Electric Medical Systems) visible light, which is then fed to photodiodes for imaging purposes. In order to achieve high-resolution images, the scintillator material is often diced into small pieces and assembled in a pixilated array with desired geometry prior to attaching to the photodiode. Reflector materials are used between the pixels to prevent light penetration from one pixel to another and to enhance light output of each unit, and thus improve imaging resolution. Light-absorbing materials are also typically added to further minimize cross talk among pixels.
Currently, CT reflector materials consist of organic resin (binders) and titanium dioxide (TiO2) fillers (reflective materials). In addition, chromium compounds are added to the matrix as absorbing materials to minimize cross talk. These are also known as cast reflectors. However, there are several deficiencies associated with these systems.
For example, systems utilizing chromium materials have relatively low light output due to the chromium based materials. Also, these systems have relatively high cross talk. Further, there is low geometric efficiency due to the inherent thickness of the reflector. Finally, there is performance degradation under X-ray radiation because of the organic matrix.
In the past, silver was investigated as a potential replacement for titanium dioxide-based reflectors. In theory, silver allows for a much thinner reflector (less than about 10 microns), thereby making a reflector with higher reflectance, minimal cross talk, and no significant degradation under X-ray radiation. However, direct application of silver coatings (i.e. sputtered silver) onto scintillator materials resulted in low light output and undesired adhesion properties. On the other hand, silver opticlad, a commercial product based on silver reflectors and described in U.S. Pat. No. 4,720,426, which is herein incorporated by reference, gave a good initial performance but failed eventually due to silver degradation over time.
There is therefore a need to develop new CT reflectors having enhanced performance.